The present invention relates generally to an imaging device, and more particularly to a system and method for mapping a detected radiation event to a particular detector in a medical imaging device such as a positron emission tomography scanner.
A positron emission tomography (PET) scanner detects gamma rays which emanate from the patient. In a PET scan, the patient is initially injected with a radiopharmaceutical, which is a radioactive substance such as FDG ([18F] fluorodeoxyglucose) which emits positrons as it decays. Once injected, the radiopharmaceutical becomes involved in certain known bodily processes such as glucose metabolism or protein synthesis, for example. The emitted positrons travel a very short distance before they encounter an electron, at which point an annihilation event occurs whereby the electron and positron are annihilated and converted into two gamma rays. Each of the gamma ray has an energy of 511 keV, and the two gamma rays are directed in nearly opposite directions. The two gamma rays are detected essentially simultaneously by two of the detector crystals (also commonly referred to as “scintillators” or “scintillator crystals”) in the PET scanner, which are arranged in rings around the patient bore. The simultaneous detection of the two gamma rays by the two detector crystals is known as a “coincidence event.” The millions of coincidence events which are detected and recorded during a PET scan are used to determine where the annihilation events occurred and to thereby reconstruct an image of the patient.
One of the challenges in designing a high resolution PET scanner relates to the space requirements of the electronics associated with the detector crystals, in particular the photomultiplier tubes (PMTs) which are situated behind the detector crystals. The function of the photomultiplier tubes is to receive photons produced by the scintillator crystals and to generate an analog signal with a magnitude representative of the number of photons received. The photomultiplier tubes typically cannot be diminished in size beyond a certain point, so that generally each photomultiplier tube is situated behind a number of smaller detector crystals. For example, a detector module in a PET scanner may comprise a 2×2 array of photomultiplier tubes situated behind a 6×6 array of scintillator crystals. The photons generated by a detector crystal generally spread out to a certain extent and travel into adjacent detector crystals such that each of the four photomultiplier tubes typically receives a certain number of photons as a result of a gamma ray hitting a single detector crystal.
In response to a scintillation event, each PMT produces an analog signal which rises sharply when a scintillation event occurs then tails off exponentially. The relative magnitudes of these analog signals are used to calculate a coordinate pair (x, z) which indicates the position at which the gamma ray was incident on the crystal detector array. The (x, z) coordinate pair is mapped to a specific detector crystal in the detector array using a look-up table. The data in the look-up table may be referred to as a detector position map. For each pair of coordinates (x, z), the detector position map maps that pair of coordinates to one of the detectors in the array of detectors.
With ideal hardware, the detector position map would simply be a rectangular grid corresponding to the geometry of the detector crystals, where each detector crystal would have an equal area on the grid. However, this simple mapping scheme is generally not accurate due to non-linearity in the analog parts of the detector and manufacturing differences between detector crystals. For example, if a simple rectangular grid were used as the detector position map, a certain percentage of (x, z) coordinate pairs would be mapped to the wrong detector crystal.
The present invention provides a system and method for producing a detector position map which accurately maps the detected coordinates (x, z) to the correct detector crystal.